Analyte system and method for determining hemoglobin parameters in whole blood

ABSTRACT

A replaceable cuvette assembly for use in an optical absorbance measurement system for measuring whole-blood hemoglobin parameters or whole-blood bilirubin parameters. The replaceable cuvette assembly includes a cuvette substrate and a cuvette module fixedly connected to the cuvette substrate wherein the cuvette substrate is a support for securing the cuvette assembly within the optical absorbance measurement system. The cuvette module has a sample inlet port, a sample outlet port, an electronic chip assembly, a sample receiving chamber that fluidly communicates with the sample inlet port and the sample outlet port, a first cuvette window, and a second cuvette window forming a portion of the sample receiving chamber. The first cuvette window and the second cuvette window are aligned with each other defining a cuvette optical path length between the first cuvette window and the second cuvette window and disposed within an optical path of the optical absorbance measurement system.

BACKGROUND OF THE INVENTION 1. Field of the Invention

The present invention relates generally to spectroscopic systems andmethods for the identification and characterization of hemoglobinparameters in blood.

2. Description of the Prior Art

An ultraviolet-visible light spectroscopic system involves absorptionspectroscopy or reflectance spectroscopy. As the name implies, suchsystems use light in the visible and near ultraviolet ranges foranalyzing a sample. The wavelength range is typically from about 400 nmto about 700 nm. The absorption or reflectance of the visible lightdirectly affects the perceived color of the chemicals involved. UV/Visspectroscopy is routinely used in analytical chemistry for thequantitative determination of different analytes, such as transitionmetal ions, highly conjugated organic compounds, and biologicalmacromolecules. Spectroscopic analysis is commonly carried out insolutions but solids and gases may also be studied.

A near-infrared spectroscopic system also involves absorptionspectroscopy or reflectance spectroscopy. Such systems use light in thenear-infrared range for analyzing a sample. The wavelength range istypically from about 700 nm to less than 2,500 nm. Typical applicationsinclude pharmaceutical, medical diagnostics (including blood sugar andpulse oximetry), food and agrochemical quality control, and combustionresearch, as well as research in functional neuroimaging, sportsmedicine & science, elite sports training, ergonomics, rehabilitation,neonatal research, brain computer interface, urology (bladdercontraction), and neurology (neurovascular coupling).

Instrumentation for near-IR (NIR) spectroscopy is similar to instrumentsfor the UV-visible and mid-IR ranges. The basic parts of aspectrophotometer are a light source, a holder for the sample, adiffraction grating in a monochromator or a prism to separate thedifferent wavelengths of light, and a detector. The radiation source isoften a Tungsten filament (300-2500 nm), a deuterium arc lamp, which iscontinuous over the ultraviolet region (190-400 nm), Xenon arc lamp,which is continuous from 160-2,000 nm, or more recently, light emittingdiodes (LED) for the visible wavelengths. The detector is typically aphotomultiplier tube, a photodiode, a photodiode array or acharge-coupled device (CCD). Single photodiode detectors andphotomultiplier tubes are used with scanning monochromators, whichfilter the light so that only light of a single wavelength reaches thedetector at one time. The scanning monochromator moves the diffractiongrating to “step-through” each wavelength so that its intensity may bemeasured as a function of wavelength. Fixed monochromators are used withCCDs and photodiode arrays. As both of these devices consist of manydetectors grouped into one or two dimensional arrays, they are able tocollect light of different wavelengths on different pixels or groups ofpixels simultaneously. Common incandescent or quartz halogen light bulbsare most often used as broadband sources of near-infrared radiation foranalytical applications. Light-emitting diodes (LEDs) are also used. Thetype of detector used depends primarily on the range of wavelengths tobe measured.

The primary application of NIR spectroscopy to the human body uses thefact that the transmission and absorption of NIR light in human bodytissues contains information about hemoglobin concentration changes. Byemploying several wavelengths and time resolved (frequency or timedomain) method and/or spatially resolved methods, blood flow, volume andabsolute tissue saturation (StO₂ or Tissue Saturation Index (TSI)) canbe quantified. Applications of oximetry by NIRS methods includeneuroscience, ergonomics, rehabilitation, brain computer interface,urology, the detection of illnesses that affect the blood circulation(e.g., peripheral vascular disease), the detection and assessment ofbreast tumors, and the optimization of training in sports medicine.

With respect to absorption spectroscopy, the Beer-Lambert law statesthat the absorbance of a solution is directly proportional to theconcentration of the absorbing species in the solution and the pathlength. Thus, for a fixed path length, UV/Vis and NIR spectroscopy canbe used to determine the concentration of the absorber in a solution.The method is most often used in a quantitative way to determineconcentrations of an absorbing species in solution, using theBeer-Lambert law:

A=log₁₀(I ₀ /I)=εcL

where A is the measured absorbance, in Absorbance Units (AU),

I₀ is the intensity of the incident light at a given wavelength,

I is the transmitted intensity,

L the path length through the sample, and

c the concentration of the absorbing species.

For each species and wavelength, ε is a constant known as the molarabsorptivity or extinction coefficient. This constant is a fundamentalmolecular property in a given solvent, at a particular temperature andpressure, and has units of 1/M*cm or often AU/M*cm. The absorbance andextinction ε are sometimes defined in terms of the natural logarithminstead of the base-10 logarithm.

The Beer-Lambert Law is useful for characterizing many compounds butdoes not hold as a universal relationship for the concentration andabsorption of all substances.

It is recognized by those skilled in the art that various factors affectthese spectroscopic systems. These factors include spectral bandwidth,wavelength error, stray light, deviations from the Beer-Lambert law, andmeasurement uncertainty sources.

Stray light is an important factor that affects spectroscopic systems.Stray light causes an instrument to report an incorrectly lowabsorbance.

Deviations from the Beer-Lambert law arise based on concentrations. Atsufficiently high concentrations, the absorption bands will saturate andshow absorption flattening. The absorption peak appears to flattenbecause close to 100% of the light is already being absorbed. Theconcentration at which this occurs depends on the particular compoundbeing measured.

Measurement uncertainty arises in quantitative chemical analysis wherethe results are additionally affected by uncertainty sources from thenature of the compounds and/or solutions that are measured. Theseinclude spectral interferences caused by absorption band overlap, fadingof the color of the absorbing species (caused by decomposition orreaction) and possible composition mismatch between the sample and thecalibration solution.

SUMMARY OF THE INVENTION

It is known that human hemoglobin (HGB) is an oxygen carrying protein inerythrocytes. The determination of its concentration in whole blood is auseful and important diagnostic tool in clinical biochemistry. COOxanalyzers are used to measure the hemoglobin parameters of blood, suchas total hemoglobin (tHb), carboxyhemoglobin (COHb), deoxyhemoglobin(HHb), oxyhemoglobin (O2Hb), methemoglobin (MetHb), and fetal hemoglobin(FHb) as well as total bilirubin (tBil) using optical absorbancemeasurements. In practice, typical COOx analyzers use lysed bloodinstead of whole blood because of the problems encountered withspectrometric analysis of whole blood. The measurement of lysed blood isrelatively straightforward since the lysing process dissolves the redblood cells and turns the blood into an almost non-diffusing medium. Theabsorbance is measured with a simple collimated beam through the cuvettewith little loss of light due to scattering. Because of the low loss oflight due to scattering, a straightforward linear analysis may be usedto find the hemoglobin and total bilirubin parameters.

Measurement of hemoglobin and total bilirubin parameters using a wholeblood sample is very challenging due to the strong optical scattering ofwhole blood. These problems are primarily related to handling theincreased light scattering level of whole blood as compared to lysedblood. This introduces light loss and nonlinear absorbance into themeasurement.

The components in a prism-based spectrometer naturally have a low straylight profile. The major contributing factor to stray light performanceis related to how the components are used.

Although the problems are primarily related to handling the increaselight scattering level of whole blood, it is not a single factor that,if resolved, is capable of solving these difficult problems. Theinventors have identified several factors that need to be addressed inorder to measure hemoglobin parameters in whole blood. Because wholeblood is a very diffuse medium, it is necessary to collect as much lightas possible to reduce the requirement for an upper absorbancemeasurement range. It is also necessary to expand the upper limit of themeasured absorbance due to the lower range of detector linearitycorrection. Blood settling effects are another problem that leads topoor correlation of absorbance of whole blood scans to absorbance oflysed blood scans. Basically, the blood cells are forming clumps orrouleaux. LED white light source brightness must also be increased.Lastly, new algorithms other than linear-based algorithms are needed toovercome the light scattering effects of whole blood.

Typical collection optics for systems using lysed blood are designed tocollect light from the cuvette in a cone of about +/−0.7 degrees wideand have an upper measure absorbance limit of 1.5 A.U. (absorbanceunits). It was discovered by the inventors that for whole blood thesystem needs to collect light from the cuvette in a cone of about +/−12degrees and that the upper absorbance limit had to increase to about 3.5A.U. As for blood settling effects, the typical time it takes to measurethe absorbance spectrum (approx. 1 minute), the whole blood in thecuvette is settling and the blood cells are forming clumps or rouleaux.Consequently, the scattering effects and the absorbance change withtime. The inventors discovered that changing the spectrometer control tocollect multiple scans frequently rather than a few scans averaged overa longer period avoided step functions in the composite absorbance scan,which is stitched together from scans from several integration times.Unfortunately, adding more scans to expand the absorbance upper limitincreases the data collection time. To resolve this dilemma, integrationtime was lowered from 5 msec to 1.2 msec to reduce data collection time.It was discovered, however, that this only works if the light level isincreased by a corresponding factor. Thus, the LED white lightbrightness must be increased.

The optical absorbance measurement of a diffuse sample such as wholeblood presents a unique problem. The diffuse transmittance of the wholeblood sample scrambles the initial spatial light distribution of themeasurement system caused by the non-uniformity typical of lightsources. Thus, the spatial light distribution of the “blank” scan can bequite different from the whole blood sample scan. Since opticaldetectors have response that varies spatially, the response can vary dueto spatial distribution changes of the incident light, even if theoverall intensity has not changed. An absorbance scan which is based onthe ratio of the whole blood sample scan to the blank scan will have asignificant absorbance component due to this this non-uniformity of thelight source in addition to the absorbance due to the sample alone. Thisresults in a significant measurement error of the whole blood sampleabsorbance that is intolerable for cooximetry.

It was discovered that, by placing the sample cuvette between diffusers,the spatial light distribution appears the same for the blank and samplescans, thus, removing this error effect. The diffusers are speciallychosen so that they diffuse a ray of incident light into the fullacceptance cone of the optical system, but not more, so that as muchlight throughput as possible may be preserved while scrambling the raycompletely across the field.

In addition, the measurement of fetal hemoglobin parameters presentsadditional problems. These include spectral acquisition times, whichmust be faster. Instead of the typical 12 seconds, it must be 5 secondsor less. The spectral acquisition time includes integration timemultiplied by the number of co-added spectra and the processing time toproduce one spectrum (full light, dark or sample) meeting all thefollowing requirements. Absolute wavelength accuracy must be less; lessthan +0.03/−0.03 nm compared to +0.1/−0.0 nm. Wavelength calibrationmaintenance (less than +0.06/−0.0 nm versus+0.1/−0.0 nm), wavelengthcalibration drift (less than 0.024 nm/° C. compared to 0.04 nm/° C.),dark current level (less than 0.06%/° C. for maximum dynamic rangeversus 0.1%/° C. of maximum dynamic range), response nonlinearity (lessthan 0.06% after correction and less than 1.2% for lowest and highest10% of dynamic range compared to 0.1% after correction and 2.0% forlowest and highest 10% of dynamic range), scattered light level (lessthan 0.02% of maximum dynamic range for fully illuminated detector arrayversus 0.1% of maximum dynamic range for fully illuminated detectorarray), thermal drift of response (intensity change maximum of 6% andtilt max of 6% over spectral range compared to intensity change maximumof 10% and tilt max of 10% over spectral range), and temperatureexcursion allowed during measurement (less than 0.5° C. compared to 2°C.) must all be less. The present invention includes these additionalfeatures for use in measuring fetal hemoglobin parameters.

In another aspect of the present invention, commercially availablecompact and low-cost spectrometers typically use diffraction gratings(reflective or transmissive) to disperse the light input. Diffractiongratings give a high degree of dispersion in a small volume, and producea relatively constant bandwidth (or resolution) vs. wavelength preferredby the typical user. Gratings, however, suffer from high stray light dueto multiple diffraction orders and also from the imperfections inherentin the lines that are etched to produce the grating surface. Thus,mass-produced but expensive master holographic gratings are typicallyemployed in applications requiring low stray light, rather than the morecommonly available replicated gratings.

The requirement for low stray light for COOx analyzers limits thepopulation of suitable grating manufacturers to the several in the worldthat produce master holographic or individually precision photoetchedgratings. This serves to make it difficult to get low-costhigh-performance gratings in quantity.

Prisms are also used to make spectrometers. Prisms have no issues withmultiple diffraction orders and their surfaces have orders of magnitudefewer imperfections than the surface of a grating. The components in aprism-based spectrometer naturally have a low stray light profile. Thus,stray light in a prism spectrometer can potentially be lower by an orderof magnitude or more compared to a grating spectrometer of otherwisesimilar design. The major contributing factor to stray light performancearises from how the components are used. There are three main sources ofstray light. These include (1) overfilling of the spectrometer numericalaperture, (2) retroreflection from the light-array detector, and (3) thefocal plane image. Light in excess of that required to fully illuminatethe numerical aperture of the spectrometer can bounce around in thespectrometer and land on the detector. In the present invention, thenumerical aperture of the optical fiber is 0.22 and the numericalaperture of the prism spectrometer is 0.1. A stop placed above theoptical fiber input restricts the light input cone from the opticalfiber to prevent excess light input. The light-array detector does notabsorb all of the light impinging upon it, but back-reflects a portion.This retroreflection must be controlled to land into an absorbingsurface or beam trap to prevent it from scattering onto the detector.Imparting a slight tilt of the light-array detector forces theretroreflection back into a harmless direction. The image of the slit onthe detector focal plane must be as sharp as possible. Any excessiveoverfill of the detector due to defocus can be a potential source ofstray light. If this light hits detector structures such as bond wires,metallization pads, etc., it can bounce back onto the sensitive surfaceof the detector.

Additionally, a prism spectrometer spreads the blue end of the spectrumout over more pixels than a diffraction grating spectrometer and, thus,the blue end of the spectrum gives a lower signal per pixel. Tocompensate for the lower signal per pixel, an LED with higher bluepower, or a cool-white LED, is used. The signal in the blue can befurther boosted by adding an inexpensive filter glass after the LED thatslightly attenuates the red end. Kopp filter glass type 4309, about 3 mmthick, is useful for this purpose. The main disadvantage of prisms isthe lower dispersive power they have compared to a grating, and thevariation of resolution with wavelength. In the present invention when aprism is used, the former disadvantage is mitigated by using a smallenough light-array detector; the latter is mitigated because theanalysis of whole blood does not require a uniformly small resolutionacross the waveband of interest.

Currently available spectrometers typically list a uniform 1 nmresolution for the blood measurement spectral region of 455-660 nm. Inthe present invention, the spectral region is expanded and covers thespectral region of 422-695 mm. Further, the resolution is selectivelychanged upward in regions where low resolution is not required (such asthe 600-695 nm region and 422-455 nm region). In the present invention,these regions have a resolution greater than 1 nm. Typically, theresolution is about 3.0 to about 3.5 nm. These ranges are used tocapture additional wavelength calibration peaks for wavelengthcalibration and fluid detection. The larger spectral region of thepresent invention requires consideration of the dispersed spectrum fromthe prism. The dispersed spectrum must be spread out over thelight-array detector and cover enough pixels to sample the spectrum at afine enough resolution but not so much as to extend outside of thedetector array. Due to the wider spectral range, the present inventionincorporates a light-array detector having 1024 pixels with an activearea length of about 8.0 mm.

A minimal-part reference design for an optical dispersion spectrometerrequires only two optical components: a light dispersion element (i.e.prism or grating) and a doublet (achromatic) lens. The prism/grating hasa reflective coating on the base. One example of an acceptable prism isa Littrow prism. The Littrow prism has a structure such that it isusable for a compact and low-cost spectrometer of the present invention.The prism material (dispersion characteristic) and the lens focal lengthare further considerations. Although other prisms and achromatic lensesmay be used, one embodiment of the present invention incorporates aSchott F5 glass prism and an 80 mm focal length lens. This particularcombination provides a dispersion length of the spectrum of about 6.48mm. This dispersion length leaves about 0.75 mm on either end of thelight-array detector available for tolerance variations and darkcorrection pixels.

Thermal drift of the spectral response must be considered. It iscritical that the spectral response of the spectrometer stays within acertain range between the full light and whole blood scans. Any changein spectrometer response will cause absorbance errors. The mainprecaution against this change is to make sure that the image of theslit overfills the pixels so that image drift due to temperature doesnot cause a reduction of light on the detector pixel. The 1:1 imaging ofthe system combined with a 200 μm diameter optical fiber overfills the125 μm tall pixels. As long as image drift is confined to less thanabout 30 μm of movement in either direction along the detector over ameasurement interval, thermal drift is not a problem. The presentinvention also contemplates various mechanisms to minimize thermal drifteffects on the spectral response. These mechanisms include insulatingthe spectrometer housing to minimize temperature changes external to thespectrometer housing, maintaining the temperature within thespectrometer housing using a temperature-controlled heat source, and/orincorporating a temperature-compensating lens mount for the achromaticlens.

The process of the present invention that transforms the electricalsignals from the spectrometer will now be discussed. First, theabsorbance is measured, which is minus the base-ten logarithm of theratio of the electrical signal received when the blood sample is in thecuvette to the electrical signal received when a clear fluid is in thecuvette. Second, the absorbance values at each wavelength are put into amapping function that maps absorbance values to the analyte levels (COOxparameters and bilirubin) in the whole blood sample. The mappingfunction and its coefficients are established by using the absorbancevalues measured for whole blood samples with known analyte values, andestablishing the relationship between these absorbance values and theknown analyte values.

The present invention achieves these and other objectives by providing acompact, low-cost COOx analyzer subsystem.

In one embodiment of the present invention, there is a system formeasuring whole-blood hemoglobin parameters that includes (a) anoptical-sample module having a light-emitting module, a replaceablecuvette assembly, and a calibrating-light module, (b) an optical fiber,(c) a spectrometer module, and (d) a processor module. Thelight-emitting module has an LED light source capable of emitting lightwhere the light is directed along an optical path. The cuvette assemblyis adjacent the light-emitting module where the cuvette assembly isadapted for receiving a whole-blood sample and has a sample receivingchamber with a first cuvette window and a second cuvette window alignedwith each other. The sample receiving chamber is disposed in the opticalpath for receiving light from the LED light source and has a definedoptical path length between the first cuvette window and the secondcuvette window along with an electronic chip capable of storing apath-length value of the sample receiving chamber. The calibrating-lightmodule has a calibrating-light source with one or more known wavelengthsof light where the calibrating-light module is capable of emitting acalibrating light into the optical path. The optical fiber has alight-receiving end and a light-emitting end. The light-receiving endoptically connects to the optical-sample module where thelight-receiving end receives the light from the optical path andconducts the light to the light-emitting end. The spectrometer modulereceives the light from the light-emitting end of the optical fiber,separates the light into a plurality of light beams where each lightbeam has a different wavelength, and converts the plurality of lightbeams into an electrical signal. The processor module (1) obtains thepath-length value of the sample receiving chamber of the replaceablecuvette from the electronic chip and (2) receives and processes theelectrical signal from the spectrometer module generated for awhole-blood sample. The path-length value of the sample chamber is usedto transform the electrical signal into an output signal useable fordisplaying and reporting hemoglobin parameter values and/or totalbilirubin parameter values for the whole-blood sample.

In another embodiment of the present invention, the light-emittingmodule includes a plurality of optical components disposed in theoptical path between the LED light source and the cuvette assembly wherethe plurality of optical components includes at least an opticaldiffuser and one or more of a collimating lens, a circular polarizer,and a focusing lens.

In a further embodiment of the present invention, the calibrating-lightmodule includes a diffuser disposed in the optical path downstream fromthe cuvette assembly but upstream from a beam splitter.

In still another aspect of the present invention, there is disclosed anoptical absorbance measurement system for whole blood. The systemincludes an optical-sample module, an optical fiber, a spectrometermodule, and a processor module. The optical-sample module includes alight-emitting module, a cuvette module, a first optical diffuser, and asecond optical diffuser. The cuvette module is positioned between thefirst optical diffuser and the second optical diffuser. The spectrometermodule receives the light from the light-emitting end of the opticalfiber, separating the light into a plurality of light beams andconverting the plurality of light beams into an electrical signal. Theprocessor module receives and processes the electrical signal from thespectrometer module generated for the whole-blood sample and transformsthe electrical signal into an output signal useable for displaying andreporting hemoglobin parameter values and/or total bilirubin parametervalues for the whole-blood sample.

In yet another embodiment, the spectrometer module includes an inputslit positioned in the optical path to receive the light emitted fromthe light-emitting end of the optical fiber and to transmit the lighttherethrough, a light dispersing element disposed in the optical pathwhere the light dispersing element receives the light transmittedthrough the input slit, separates the light into the plurality of lightbeams where each light beam has a different wavelength, and re-directsthe plurality of light beams back toward but offset from the input slit,and a light-array detector capable of receiving the plurality of lightbeams and converting the plurality of light beams into an electricalsignal for further processing.

In another embodiment, the spectrometer module has athermal-compensating means for maintaining a position of the pluralityof light beams on the light-array detector. The thermal-compensatingmeans includes one or more of insulation disposed around thespectrometer housing, a temperature controller assembly disposed on thespectrometer housing (the temperature controller assembly being, forexample, a heating tape with a thermistor or other temperature measuringcomponent and a program that controls the heating of the tape based onthe temperature within the spectrometer housing), and athermal-compensating lens mount.

In a further embodiment, the thermal-compensating lens mount has a fixedmount end and an unfixed mount end that permits thermal expansion andcontraction of the thermal-compensating lens mount. The fixed mount endis fixedly attached to a baseplate or a bottom of the spectrometerhousing. The lens mount has a coefficient of expansion greater than thecoefficient of expansion of the baseplate or the spectrometer housing towhich the lens mount is attached. The thermal-compensating lens mountmoves linearly and transversely relative to an optical path of the lightfrom the light input slit based on the coefficient of expansion of thelens mount. This temperature-based movement of the lens mount maintainsthe position of the dispersed light from the light dispersing elementonto the light-array detector. In other words, thermal re-positioning ofthe achromatic lens by way of the thermal-compensating lens mount causesthe dispersed light from the light dispersing element to impinge ontothe light-array detector without affecting the electric signal generatedby the light-array detector from the impinging light. The shift of thelight beam is caused by the light-dispersing element reacting to atemperature change.

In another embodiment, there is disclosed a compact spectrometer formeasuring hemoglobin parameters in whole blood. The spectrometerincludes an enclosed housing having a light input end/an optical fiberhousing end with a light entrance port, a light input slit disposed onan electronic circuit substrate, the electronic circuit substratedisposed in the enclosed housing where the light input slit is alignedwith and adjacent to the light entrance port, a light-array detectordisposed on the circuit board substrate adjacent the light input slit,and an optical component group consisting of a light dispersing elementdisposed downstream from the light input slit and a spherical achromaticlens disposed between the light input slit and the light dispersingelement where the light dispersing element has a reflective surface on aback side to reflect the dispersed light back toward the achromaticlens. The achromatic lens transmits light from the light input slit tothe light dispersing element and transmits dispersed light reflectedfrom the light dispersing element to the light-array detector. Toaccomplish this, the achromatic lens is slightly off axis relative tothe light coming from the light input slit so that the dispersed lightfrom the light dispersing element is not directed back to the lightinput slit but to the light-array detector.

In a further embodiment, there is disclosed a method of measuringwhole-blood hemoglobin parameters despite strong optical scatteringcaused by whole blood. The method includes providing a light source suchas a LED light source with a spectral range of about 422 nm to about 695nm, guiding light having the spectral range from the light source alongan optical path, providing a cuvette module with a sample receivingchamber having a first cuvette window disposed in the optical path wherethe first cuvette window transmits the light through the samplereceiving chamber and through a second cuvette window aligned with thefirst cuvette window where the sample receiving chamber contains asample of whole blood, providing a pair of diffusers (i.e. a firstdiffuser and a second diffuser) disposed in the optical path where thefirst cuvette window and the second cuvette window of the samplereceiving chamber of the cuvette are disposed between the pair ofdiffusers, guiding light from the cuvette module into a spectrometerhaving a light dispersing element that separates the light into aplurality of light beams where each light beam has a differentwavelength and converts the plurality of light beams into an electricalsignal, and processing the electrical signal into an output signaluseable for displaying and reporting hemoglobin parameter values and/ortotal bilirubin parameter values of the sample of whole blood.

In another embodiment of the method, the processing step includesprocessing the electrical signal to spectral absorbance and then mappingthe spectral absorbance to hemoglobin parameter values and/or bilirubinparameter values using a computational mapping function.

In still another embodiment of the method, the processing step includesusing a kernel-based orthogonal projection to latent structures mappingfunction as the computational mapping function.

In another embodiment of the method, there is disclosed a method ofmeasuring hemoglobin parameters in a whole blood sample. The methodincludes (1) measuring and recording a transmitted light intensity scanover a plurality of wavelengths in a measurement range by transmittinglight through a cuvette module having an optical path with a knownoptical path length therethrough where the cuvette module is filled witha transparent fluid, (2) measuring and recording a transmitted lightintensity scan over the plurality of wavelengths of the measurementrange by transmitting light through the cuvette a second time having theoptical path with the known optical path length therethrough where thecuvette module is filled with a whole blood sample, wherein eachmeasuring and recording step of the transparent fluid and the wholeblood sample includes diffusing and circularly polarizing thetransmitted light before transmitting the transmitted light through thecuvette module and then diffusing the transmitted light emitting fromthe cuvette module before determining a spectral absorbance, (3)determining a spectral absorbance at each wavelength of the plurality ofwavelengths of the measurement range based on a ratio of the transmittedlight intensity scan of the whole blood sample to the transmitted lightintensity scan of the transparent fluid using a prism-basedspectrometer, and (4) correlating the absorbance at each wavelength ofthe plurality of wavelengths of the measurement range to hemoglobinparameter values and/or bilirubin parameter values of the blood sampleusing a computational mapping function.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a simplified, perspective view of one embodiment of thepresent invention showing a compact COOx subsystem.

FIG. 2 is a side elevation view of one embodiment of an optical-samplemodule shown in FIG. 1.

FIG. 3 is a front, perspective view of one embodiment of alight-emitting module of the optical-sample module shown in FIG. 2.

FIG. 3A is a front, perspective view of the light-emitting module shownin FIG. 3 showing a plurality of optical components.

FIG. 3B is an enlarged, side elevation view of the optical componentsshown in FIG. 3A.

FIG. 4 is a front perspective view of one embodiment of a cuvetteassembly of the optical-sample module shown in FIG. 1.

FIG. 5 is a rear perspective view of the cuvette assembly shown in FIG.4.

FIG. 6 is a front elevation view of a cuvette module of the cuvetteassembly showing fluid input and output ports, a sample receivingchamber, a sample window, and an electronic chip assembly.

FIG. 7 is a rear perspective view of the sample receiving chamber ofFIG. 6 showing cuvette first and second windows.

FIG. 8 is a rear plan view of the sample receiving chamber showing theelectronic chip assembly disposed adjacent the sample receiving chamber.

FIG. 9 is a perspective view of one embodiment of a calibrating lightmodule of the optical-sample module of FIG. 1.

FIG. 10 is a side cross-sectional view of the calibrating light moduleof FIG. 8 showing a calibrating light source.

FIG. 11 is a simplified, side plan view of the calibrating light sourceof the calibrating light module of FIG. 9 showing a plurality of opticalcomponents.

FIG. 12 is a front perspective view of one embodiment of a spectrometermodule of FIG. 1 with a cover removed showing the internal components.

FIG. 13 is a rear perspective view of the spectrometer module of FIG. 12showing an input light slit and adjacent light-array detector.

FIG. 14 is a rear cross-sectional view of the spectrometer module ofFIG. 12 showing a single circuit board and the location of the inputlight slit and the light-array detector.

FIG. 15 is a top view of the spectrometer module of FIG. 12 showing theoptical components with superimposed ray trace.

FIG. 16 is a ray trace showing the input light from the input light slitand a plurality of light beams refracted onto the light-array detector.

FIG. 17A is a perspective view of one embodiment of athermal-compensating means for the spectrometer module showinginsulation wrapped around the spectrometer module.

FIG. 17B is a perspective view of another embodiment of athermal-compensating means for the spectrometer module showing atemperature controlling assembly.

FIG. 17C is a cross-sectional view of one embodiment of a lens mount ofthe spectrometer module of FIG. 12 showing a temperature-compensatinglens mount.

FIG. 18 is a cross-sectional view of one embodiment of a lens mount ofthe spectrometer module of FIG. 12 showing a fixed lens mount.

FIG. 19 is a graphic illustration showing the correlation results of theCOOx analyzer subsystem of the present invention for total hemoglobinusing a K-OPLS mapping function and method.

FIG. 20 is a graphic illustration showing the correlation results of theCOOx analyzer subsystem of the present invention for oxyhemoglobin usinga K-OPLS mapping function and method.

FIG. 21 is a graphic illustration showing the correlation results of theCOOx analyzer subsystem of the present invention for carboxyhemoglobinusing a K-OPLS mapping function and method.

FIG. 22 is a graphic illustration showing the correlation results of theCOOx analyzer subsystem of the present invention for deoxyhemoglobinusing a K-OPLS mapping function and method.

FIG. 23 is a graphic illustration showing the correlation results of theCOOx analyzer subsystem of the present invention for methemoglobin usinga K-OPLS mapping function and method.

FIG. 24 is a graphic illustration showing the correlation results of theCOOx analyzer subsystem of the present invention for total bilirubinusing a K-OPLS mapping function and method.

DETAILED DESCRIPTION

Embodiments of the present invention are illustrated in FIGS. 1-24. FIG.1 shows one embodiment of a COOx analyzer subsystem 10. COOx analyzersubsystem 10 includes at least an optical-sample module 20, an opticalfiber 90 and a spectrometer module 100. COOx analyzer subsystem 10 mayoptionally include a processor module 150 or processor module 150 mayoptionally be included in an electronics circuit of a diagnostic systemin which the COOx analyzer subsystem 10 is a part. Line 5 is included tosignify that the processor module 150 may or may not be part of the COOxsubsystem 10. Processor module 150 includes, but is not limited to amicroprocessor module 152 and a memory module 154. Optionally, theprocessor module 150 may also include a converter module 156 orconverter module 156 may be external to the COOx analyzer subsystem 10.COOx analyzer subsystem 10 is used to measure the hemoglobin parametersof blood such as total hemoglobin (tHb), carboxyhemoglobin (COHb),deoxyhemoglobin (HHb), oxyhemoglobin (O2Hb), methemoglobin (MetHb), andfetal hemoglobin (FHb) as well as total bilirubin (tBil) using opticalabsorbance.

FIG. 2 illustrates optical-sample module 20. Optical-sample module 20includes a light-emitting module 22, a cuvette assembly 40 and acalibrating-light module 60. Light-emitting module 22, as the termimplies, emits a visible light beam toward the cuvette assembly 40 thatis then received by the calibrating-light module 60, which is thentransmitted to spectrometer module 100. The light beam 12 defines anoptical path 21.

FIGS. 3-3A illustrate perspective views of the embodiment oflight-emitting module 22 of FIG. 2. Light-emitting module 22 includes alight-emitting module substrate 24 that contains an electrical circuit(not shown) and a light-emitting optics assembly 25. Light-emittingoptics assembly 25 has an optics assembly housing 26 with an opticsassembly end 26 a. A beam of visible light 28 a emits from opticsassembly end 26 a of light-emitting optics assembly 25 whenlight-emitting module 22 is powered on by a signal received fromprocessor module 150. FIG. 3A illustrates light-emitting optics assembly25 with optics assembly housing 26 removed exposing a plurality ofoptical components B contained within light-emitting assembly 25.

Turning now to FIG. 3B, there is illustrated an enlarged side view ofthe plurality of optical components B of FIG. 3A. In this embodiment,optical components B includes a light-emitting diode (LED) light source28, a collimating lens 30, a first diffuser 32, a circular polarizer 34,a focusing lens 36, and an optional protective window 38. Circularpolarizer 34 provides a distinct advantage. This advantage providesimproved sensitivity and accuracy of the system. Hemoglobin has opticalrotary characteristics, which means that the polarization sensitivity ofa spectrometer will cause an absorbance error if non-circularlypolarized light is used to measure hemoglobin absorbance. Unlike forother polarization states of light, the polarization state of thecircularly polarized light is not changed when passing throughhemoglobin. Thus, the polarization response of the spectrometer is thesame for the circularly polarized light passing through the hemoglobinas it is for the reference scan taken with the cuvette filled with atransparent fluid.

FIGS. 4 and 5 illustrated front and rear perspective views of oneembodiment of the cuvette assembly 40. Cuvette assembly 40 includes acuvette substrate 41 and a cuvette module 43. Cuvette substrate 41provides a support for securing the cuvette assembly 40 within theanalyte subsystem 10 and includes a cuvette light path opening 42 thatis disposed within optical path 21 and is aligned with the light beamemitted from light-emitting module 22. Cuvette module 43 includes acuvette first portion 44 having a sample receiving recess 45, a sampleinlet port 46, a sample outlet port 47, an electronic chip assembly 48,and a first cuvette window 49, and a cuvette second portion 50 having asecond cuvette window 52 (shown in FIG. 6 and delineated as outline 53)opposite and aligned with the first cuvette window 49 where the firstand second cuvette windows 49, 52 are aligned with and dispersed withinoptical path 21. Cuvette first portion 44 and cuvette second portion 50are bonded to each with or without a gasket disposed between cuvettefirst and second portions 44, 50. Bonding may be achieved usingadhesives, ultrasonic techniques, solvent based techniques, etc. Whenassembled and as shown in FIG. 6, sample receiving recess 45 of cuvettefirst portion 44 forms a sample receiving chamber 54 with cuvette secondportion 50 that fluidly communicates with sample inlet and outlet ports46, 47. The distance between first and second cuvette windows 49, 52 ofsample receiving chamber 54 define a cuvette optical path length, whichis accurately measured and stored within electronic chip 48 for laterretrieval by processor module 150. A typical optical path length used inthis embodiment of the present invention is 0.0035 inches (0.090 mm).

Turning now to FIG. 7, there is illustrates an enlarged, rearperspective view of cuvette first and second portions 44, 50. As shown,cuvette first portion 44 has sample chamber recess 45 with first cuvettewindow 49 and electronic chip recess 48 a for receiving electronic chipassembly 48. Cuvette second portion 50 has second cuvette window 52 thatforms sample receiving chamber 54 when assembled together with cuvettefirst portion 44. Second cuvette window 52 as delineated by an outline53 on cuvette second portion 50 is a raised surface that forms awater-tight seal around sample chamber recess 45 and sample receivingchamber 54. Optionally, a thin gasket may be positioned between cuvettefirst and second portions 44, 50 to more easily ensure a water-tightseal. FIG. 8 shows a rear view of cuvette first portion 44 withelectronic chip assembly 48 disposed within electronic chip recess 48 a.Electronic chip assembly 48 includes a chip circuit board 48 b and anelectronic chip 48 c that stores the cuvette optical path length valuefor the particular cuvette module 43. First cuvette window 49 isdisposed within the optical path 21 and transmits the light beam passingthrough the sample to the calibrating light module 60, which then passesthe light beam to the spectrometer module 100.

Turning now to FIG. 9, there is illustrated one embodiment of thecalibrating light module 60. Calibrating light module 60 includes acalibrating module housing 62, a light beam receiving portion 64, acalibrating light portion 70, and an optic fiber portion 80 wherecalibrating module housing 62, light beam receiving portion 64 and opticfiber portion 80 are aligned with optical path 21. Calibrating lightportion 70 is spaced from and transverse to optical path 21.

FIG. 10 is a cross-sectional, elevation view of calibrating light module60. Calibrating module housing 62 includes a first tubular conduit 62 abetween a light beam input opening 62 b and a light beam exit opening 62c as well as a second tubular conduit 62 d that is transverse to andintersects with first tubular conduit 62 a on one end and has acalibrating light beam opening 62 e on an opposite end.

Light beam receiving portion 64 houses a collimating lens 66 thatcollimates light beam 28 a received along optical path 21 from cuvettemodule 43 and directs light beam 28 a into first tubular conduit 62 a.Disposed within calibrating module housing 62 is beam splitter holderassembly 67 that is disposed transversely across first tubular conduit62 a. Beam splitter holder assembly 67 has an upward slanting surface 67a facing calibrating light beam opening 62 e and light beam exit opening62 c within optical path 21. Beam splitter holder assembly 67 supports asecond diffuser 68 and a beam splitter 69 (shown in FIG. 11) that isdisposed downstream along optical path 21 from second diffuser 68 sothat it is positioned to receive calibrating light beam 72 a and directit along optical path 21 and first tubular conduit 62 a to light beamexit opening 62 c.

Calibrating light portion 70 includes a calibrating light source 72disposed adjacent but spaced from optical path 21 that is capable ofdirecting a calibrating light beam 72 a into calibrating module housing62 through a calibrating light opening 62 e transversely to optical path21 toward beam splitter holder assembly 67. Within calibrating lightportion 70, there is a collimating lens 74 that collimates calibratinglight beam 72 a before it is reflected by beam splitter assembly 67toward light beam exit opening 62 c.

Optic fiber portion 80 is located within optical path 21 at or in thevicinity of light beam exit opening 62 c. Optic fiber portion 80includes a focusing lens 82 and a optic fiber connector assembly 84 thatincludes a connector housing 86 adapted for receiving an optical fiberassembly 90. Optic fiber portion 80 is adapted to insure that light beam28 a is properly focused by focusing lens 82 into optical fiber assembly90.

FIG. 11 is a simplified illustration of FIG. 10 showing the positionalrelationship of the optical components 66, 68, 69, 74, 82 and lightbeams 28 a, 72 a as well as optical fiber assembly 90. As can be seenfrom FIG. 11, light beam 28 a is received by collimating lens 66,transmitted through second diffuser 68 and beam splitter 69 to focusinglens 82 and into optical fiber assembly 90. As previously discussed, theimportance of using a pair of diffusers (first diffuser 32 and seconddiffuser 68) with cuvette module 43 in between the pair of diffusers 32,68 is that the spatial light distribution will appear the same for theblank scan and the whole blood sample scan. The use of diffusers 32, 68in this arrangement removes the error effect caused by nonuniformity ofthe light source and/or variation in the spatial distribution changes ofthe incident light even if the overall intensity has not changed.Diffusers 32, 68 are chosen so that they diffuse a ray of incident lightinto the full acceptance cone of the optical component group 120 of thespectrometer module 100. This effectively scrambles the ray completelyacross the optical measuring field.

Calibrating light beam 72 a when activated is received by collimatinglens 74, transmitted to beam splitter 69 and directed to focusing lens82 where it is focused into optical fiber assembly 90. Calibrating lightbeam 72 a has specific wavelengths of light used for calibrating thewavelength scale of spectrometer module 100. One example of anacceptable calibrating light source 72 is a krypton (Kr) gas dischargelamp, which provides seven Kr line wavelengths in nanometers coveringthe range of 422 to 695 nm. Prism 131 of light dispersion component 130has a nonlinear dispersion versus wavelength that requires a polynomialor other function of a higher order. The present invention uses a 5thorder polynomial to the pixel locations of the Kr line peaks to provideresidual errors well below the absolute wavelength accuracy requirementof +/−0.03 nm.

Optical fiber assembly 90 includes an optical fiber 92, a first opticalfiber connector 94 and a second optical fiber connector 96 (shown inFIG. 12). First optical fiber connector 94 is secured to a lightreceiving end 92 a of optical fiber 92 and directly and removablyconnects to connector housing 86 of optic fiber connector assembly 84.One embodiment of optical fiber 92 includes a 200 μm silica core fiberwith a numerical aperture (NA) of 0.22.

Turning now to FIGS. 12 and 13, there is illustrated one embodiment ofspectrometer module 100. Spectrometer module 100 includes a spectrometerhousing 102, a spectrometer base 104, a spectrometer cover 106 (shown inFIG. 1), an optical fiber housing end 108, and an electrical signaloutput coupler 103. Spectrometer module 100 has an outside envelopedimension of 11 cm×8 cm×2 cm and optionally includes thermalcompensation structures discussed later. Within spectrometer housing 102are contained the essential components of spectrometer module 100. Thesecomponents include a light-receiving and converting assembly 110 and anoptical component group 120. Optical component group 120 includes anachromatic lens assembly 121 and a light dispersing element 130. Lightdispersing element 130 may be a prism 131 or a grating 136. Opticalfiber assembly 90 is removably secured to optical fiber housing end 108at light entrance port 109, which optical fiber assembly 90 transmitsthe light beams 28 a, 72 a to spectrometer module 100. As previouslymentioned, light beam 28 a represents the light transmitted fromlight-emitting module 22 through cuvette module 43 whereas light beam 72a is the calibrating light transmitted from calibrating light module 60,which is used to calibrate spectrometer module 100.

Achromatic lens assembly 121 includes a lens mount 122 and a sphericalachromatic lens 124. Achromatic lens 124 receives light beams 28 a, 72a, as the case may be, and directs the light beam to light dispersionelement 130, which in this embodiment is prism 131. Prism 131 has areflective coating 132 on an outside back surface. Prism 130 refractslight beam 28 a and reflects the light back through achromatic lens 124.

Light-receiving and converting assembly 110 is securely mounted adjacentan inside surface 108 a of optical fiber housing end 108.Light-receiving and converting assembly 110 includes a circuit boardsubstrate 112 upon which is mounted a light input slit 114 that isaligned with light-emitting end 92 b (not shown) of optical fiber 92.Adjacent input slit 114 is a light-array detector 116 that receives therefracted light from prism 131. Light-array detector 116 converts therefracted light to an electrical signal, which is output through outputconnector 118 to processor module 150. Providing light input slit 114and light-array detector 116 adjacent each other on circuit board 112has several advantages. This feature greatly simplifies the constructionand improves the precision of spectrometer module 100. Otherspectrometers place these items on separate planes, where they haveseparate mounting structures, and have to be adjusted independently.This feature of mounting the input slit and light-array detectoradjacent each other on circuit board 112 eliminates the need to mountand position each structure (i.e. slit and detector) separately.

FIG. 14 is an enlarged view of light-receiving and converting assembly110. Light input slit 114 is 15 μm wide by 1000 μm long that projects anoptical fiber-slit image that is a rectangle approximately 15 μm wide by200 μm high onto the light-array detector 116 (Hamamatsu S10226-10 is anexample of a usable light array detector). Input slit 114 is applieddirectly onto the same circuit board substrate 112 as and in closeproximity to light-array detector 116. Light-array detector 116 has apixel height between about 100 to about 150 μm, which allows aone-to-one imaging of the 200 μm diameter optical fiber onto thedetector. In this embodiment, input slit 114 is laser etched in aprecise position relative to light-array detector 116 making alignmentless labor intensive. Because input slit 114 and light-array detector116 are only slightly off-axis relative to the center axis of theachromatic lens 124, there is minimal aberration and a one-to-oneimaging on light-array detector 116 is possible so that no cylindricalfocusing lens is required to shrink the optical fiber image (200 μmdiameter fiber) to match the pixel height of light-array detector 116.

Turning now to FIG. 15, there is a top view of spectrometer module 100of FIG. 13. Superimposed onto FIG. 15 is a ray trace diagram 140 of thelight beam delivered to spectrometer module 100 by optical fiber 92. Asshown, light beam 28 a enters spectrometer module 100 through input slit114 toward achromatic lens 124. Achromatic lens 124 is used off-axis;that is, the achromatic lens is slightly off-axis to the light beam 28a. Light beam 28 a is transmitted by achromatic lens 124 to prism 131,where light beam 28 a is refracted into a plurality of light beams 138a, 138 b, 138 c of different wavelengths as prisms are ought to do. Theplurality of light beams 138 a, 138 b, 138 c are reflected by prism 131back through achromatic lens 124. Achromatic lens 124 is used off-axisin order to direct the plurality of refracted and reflected light beams138 a, 138 b, 138 c from prism 131 onto light-array detector 116.

FIG. 16 is an enlarged view of ray trace diagram 140. Achromatic lens124 is used off-axis relative to entering light beam 28 a. By usingachromatic lens 124 off-axis along with prism 131 having a reflectivecoating 132 on a base of prism 131, there is achieved a compact,simplified, minimal-component spectrometer module 100 capable of beingused for measuring hemoglobin parameters and/or total bilirubinparameters in whole blood.

A change in temperature has a greater effect on beam refraction anglewhen using a prism instead of a diffraction grating. In the presentinvention, a thermal-compensating means 160 is provided to compensatefor a thermal shift in the incoming light beam caused by thelight-dispersing element 130. A temperature change within spectrometermodule 100 causes a thermally-induced movement of the slit image frominput slit 114 on light-array detector 116 caused in turn bythermally-induced changes in refractive index of the dispersive prism131. FIG. 16 shows the direction of movement of the image on light-arraydetector 116 for the thermal refractive index change in prism 131 witharrow 400. If the lens 124 is moved in the opposite direction over thesame temperature interval as indicated by arrow 402, the slit image willbe moved back to where it should be onto light-array detector 116. Toprevent this shift, the thermal-compensating means 160 may be a simpleas wrapping spectrometer module 100 with insulation to minimizetemperature change within spectrometer module 100 from a temperaturechange occurring outside of spectrometer module 100 or to placespectrometer module 100 within a temperature controlled space. Anothermeans is to include a temperature controller assembly 170 that includesat least a ribbon heater 172 attached to an inside surface or an outsidesurface of the spectrometer housing 102 and a temperature sensor 174such as thermocouple or thermistor to measure the temperature of thespectrometer housing and a heater circuit to maintain a predefinedconstant temperature. FIGS. 17A and 17B illustrate these possibilities.

In one embodiment shown in FIG. 17C, achromatic lens mount 122 is athermal-compensating lens mount. Thermal-compensating lens mount 122 hasa fixed mount end 122 a and an unfixed mount end 122 b. Fixed mount end122 a is fixedly secured to spectrometer base 104 or a baseplate 104 athat is securely attached to spectrometer base 104. Unfixed mount end122 b typically has a fastener 126 that extends through a lens mountslot 122 c of lens mount 122 and into spectrometer base 104 or baseplate104 a. Between a head 126 a of fastener 126 and lens mount 122 is ahold-down spring 128. There is sufficient spacing between lens mountslot 122 c and fastener 126 to permit expansion/contraction of lensmount 122 caused by a temperature change. The coefficient of expansionof lens mount 122 is greater than the coefficient of expansion ofspectrometer base 104 and/or baseplate 104 a so that unfixed mount end122 b permits thermal expansion and contraction of thermal-compensatinglens mount 122 in a direction shown by arrow 500, which is linear andtransverse to the light beam from input slit 114. This structure allowsachromatic lens 124 to slide relative to other components mounted onbaseplate 104 a and/or spectrometer base 104. Thermal-compensated lensmount 122 ensures that the plurality of light beams 138 a, 138 b, 138 cwill always impinge with sufficient intensity onto light-array detector116 without affecting the electrical signal generated by light-arraydetector 116 notwithstanding a temperature change within spectrometerhousing 102. One such material that meets the requirement that lensmount 122 have a greater coefficient of expansion than spectrometer base104 and/or baseplate 104 a (as the case may be) is a plastic that is amodified polyphenylene ether (PPE) resin consisting of amorphous blendsof polyphenylene oxide (PPO) polyphenylene ether (PPE) resin andpolystyrene sold under the trademark NORYL®.

FIG. 18 illustrates an alternative embodiment of lens mount 122. In thisembodiment, lens mount 122 has two fixed mount ends 122 a, where eachend 122 a is secured to baseplate 104 a and/or spectrometer base 104 byfastener 126. Because both ends 122 a of lens mount 122 are fixed, anytemperature change within spectrometer module 100 will affect angle ofthe plurality of light beams 138 a, 138 b, 138 c and where they impingeon light-array detector 116. As previously disclosed regarding the slitimage and the length of the light-array detector 116, a temperaturechange of greater than 0.5° C. will cause the intensity of one of thelight beams to not impinge completely on the light-array detectorthereby causing an inaccurate reading. To nullify this potential effect,spectrometer module 100 is equipped with a temperature controllerassembly (not shown) so that prism 131 and achromatic lens assembly 121remain at a constant temperature. Although there are several methodsavailable for maintaining the inside of spectrometer module 100 at aconstant temperature, one example of such a temperature controllerassembly to accomplish this is a ribbon heater with a thermistor (notshown) adhesively attached to the inside or outside of spectrometermodule 100, which ribbon heater is controlled by an electronicregulation circuit (not shown). Optionally, spectrometer module 100 mayalso be insulated either inside or outside or both to more easilymaintain a given temperature and protect against changes in temperaturein the vicinity surrounding spectrometer module 100. Other mechanismsinclude placement of spectrometer module 100 within a temperaturecontrolled environment.

Learning Data:

A data set of about 180 blood samples from approximately 15 differentindividuals was developed. The blood samples were manipulated usingsodium nitrite to raise MetHb values, and using CO gas to raise COHbvalues. Plasma was removed from or added to samples to change the tHblevel. Bilirubin spiking solution was added to vary the tBil level. Atonometer was used to manipulate the oxygen level. The blood sampleswere manipulated to cover a large range of analyte values. The bloodsamples were then measured on a reference lysing pHOx Ultra analyzerequipped with COOx analyzer and analysis software. The whole bloodspectra were gathered on a pHOx Ultra analyzer equipped with thehigh-angle collection optics and other modifications of the presentinvention, as described earlier, with the lyse supply line completelydisconnected and the whole blood samples running directly into thecuvette assembly 40 without lyse or any other dilution. Both analyzerswere equipped with Zeonex windows in the respective cuvettes. This dataset has been turned into a Matlab cell array file for use with Matlabscripts.

Prediction Model:

The next step in the calculation is to create a prediction model. Threemodels were developed for the analysis: one for the COOx parameters tHband COHb, a second for HHb and MetHb, and a third for tBil. The quantityfor O2Hb was determined by subtracting COHb, HHb, and MetHb from 100%.The X-data array was constructed from terms created from the measuredabsorbance at the wavelengths between 462-650 nm, 1 nm spacing. The tBilmodel was developed using the same set of data as the COOx model, exceptthat samples with MetHb values greater than or equal to 20% were leftout of the model. For each model, five Y-predictive values were assigned(O2Hb, HHb, COHb, MetHb, tBil) with tHb determined by adding the resultsfor O2Hb, HHb, COHb, and MetHb. The number of Y-orthogonal values neededwas determined by manual optimization of the correlation residual of themapping function blood predictions with the reference analyzer values.

Using an initial calibration data set, the calibration sequence of amachine learning algorithm establishes a relationship between a matrixof known sample characteristics (the Y matrix) and a matrix of measuredabsorbance values at several wavelengths and potentially other measuredvalues based on absorbance versus wavelength (the X matrix). Once thisrelationship is established, it is used by the analyzer to predict theunknown Y values from new measurements of X on whole blood samples.

Table 1 summarizes the settings and inputs used for the optimizedmodels. The X-data consists of the absorbance and other terms based onabsorbance vs. wavelength. In the process of optimizing the model,absorbance derivatives vs. wavelength were added. Models for analytesmore sensitive to nonlinear scatter effects were built up with squareroot terms of the absorbance and its derivative. The model for analytesmore affected by scatter had a correction term proportional to thefourth power of the wavelength. The X-vector row has one value for eachwavelength for each of the three absorbance-based terms f, g, and hshown in the table for each model.

TABLE 1 Parameters used to construct algorithm models (KOPLS method).Kernel Y-predictive Y-orthogonal X data structure polynomial Modelcomponents components (from absorbance vs. wavelength) exponent tHb,COHb 5  4${{f(\lambda)} = \sqrt{\frac{{dA}(\lambda)}{d\; \lambda}}},{{g(\lambda)} = \frac{{dA}(\lambda)}{d\; \lambda}},{{h(\lambda)} = \sqrt{A(\lambda)}}$0.5 HHb, MetHb 5  4${{f(\lambda)} = \frac{{dA}(\lambda)}{d\; \lambda}},{{g(\lambda)} = {{A(\lambda)} \cdot \left( \frac{\lambda}{650\mspace{14mu} {nm}} \right)^{4}}},{{h(\lambda)} = {A(\lambda)}}$1.0 tBil 5 16${{f(\lambda)} = \sqrt{\frac{{dA}(\lambda)}{d\; \lambda}}},{{g(\lambda)} = \sqrt{A(\lambda)}},{{h(\lambda)} = {A(\lambda)}}$1.0

The calibration set Y matrix is built up as follows from the knownvalues of the calibration sample set of n lysed blood samples:

$Y = \begin{bmatrix}{tHb}_{1} & {COHb}_{1} & {HHHb}_{1} & {MetHb}_{1} & {tBil}_{1} \\{tHb}_{2} & {COHb}_{2} & {HHHb}_{2} & {MetHb}_{2} & {tBil}_{2} \\{\ldots\ldots} & {\ldots\ldots} & {\ldots\ldots} & {\ldots\ldots} & {\ldots\ldots} \\{tHb}_{n} & {COHb}_{n} & {HHHb}_{n} & {MetHb}_{n} & {tBil}_{n}\end{bmatrix}$

where tHb is the total hemoglobin value of the lysed blood sample,

-   -   COHb is the carboxyhemoblogin value of the lysed blood sample,    -   HHb is the deoxyhemoglobin value of the lysed blood sample,    -   MetHb is the methemoglobin value of the lysed blood sample, and    -   tBil is the total bilirubin value of the lysed blood sample.

The X matrix is structured as follows:

$X = \begin{bmatrix}{{f_{1}\left( \lambda_{1} \right)},{\ldots \; {f_{1}\left( \lambda_{n} \right)}},} & {{_{1}\left( \lambda_{1} \right)},{{\ldots }_{n}\left( \lambda_{n} \right)},} & {{h_{1}\left( \lambda_{1} \right)},{\ldots \; {h_{1}\left( \lambda_{n} \right)}}} \\\vdots & \ddots & \vdots \\{{f_{n}\left( \lambda_{1} \right)},{\ldots \; {f_{n}\left( \lambda_{n} \right)}},} & {{_{n}\left( \lambda_{1} \right)},{{\ldots }_{n}\left( \lambda_{n} \right)},} & {{h_{n}\left( \lambda_{1} \right)},{\ldots \; {h_{n}\left( \lambda_{n} \right)}}}\end{bmatrix}$

where: f,g,h are the absorbance-based functions listed in Table 1 versuswavelength, respectively.

The matrix X includes contributions from absorbance at the variouswavelengths. The scope of the invention includes optionally adding othermeasurements to the calculation to reduce interferent effects.

Once these matrices are formed, they are used as the calibration set andthe mapping function is computed according to the procedures particularto the machine learning algorithm chosen.

As described previously, conventional partial least squares, linearregression, linear algebra, neural networks, multivariate adaptiveregression splines, projection to latent structures, kernel-basedorthogonal projection to latent structures, or other machine learningmathematics is used with results obtained from the calibration set ofdata to determine the empirical relationship (or mapping function)between the absorbance values and the hemoglobin parameters. Typically,a mathematics package is used to generate the results where the packagegenerally has options to select one of the machine learning mathematicsknown to those skilled in the art. Various mathematics packages existand include, but are not limited to, Matlab by MatWorks of Natick,Mass., “R” by R Project for Statistical Computing available over theInternet at www.r-project.org, Python from Python Software Foundationand available over the Internet at www.python.org in combination withOrange data mining software from Orange Bioinformatics available overthe Internet at orange.biolab.si, to name a few.

It will be shown that the method of Kernel-Based Orthogonal Projectionto Latent Structures (KOPLS) may be used as one type of machine learningalgorithm to generate the mapping function. An explanation anddescription of KOPLS is best exemplified by the following references:Johan Trygg and Svante Wold. “Orthogonal projections to latentstructures (O-PLS).” J. Chemometrics 2002; 16: 119-128; MattiasRantalainen et al. “Kernel-based orthogonal projections to latentstructures (K-OPLS).” J. Chemometrics 2007; 21: 376-385; and Max Bylesjoet al. “K-OPLS package: Kernel-based orthogonal projections to latentstructures for prediction and interpretation in feature space.” BMCBioinformatics 2008, 9:106, which references are incorporated herein byreference. The kernel-based mathematics is useful in handling non-linearbehavior in systems by using a kernel function to map the original datato a higher order space. Although any of the previously describedmachine learning mathematics may be used to enable one of ordinary skillin the art to practice the present invention, KOPLS has an additionaladvantage over other calculations such as, for example, conventionalpartial least squares because it can not only establish a relationshipbetween quantified variations and analyte values to be determined, butcan also remove unquantitated yet consistently present variation in theoriginal data. These unquantitated variations might be due to analyzerand/or blood effects such as scatter losses and other interferingphenomena that are not explicitly measured. By extracting theseunquantitated variations from the data, the method leaves behind in thedata the information used to predict the measured values.

Using an initial training data set, the KOPLS model establishes arelationship (mapping function) between the matrix of known samplecharacteristics (the H matrix), and a matrix of measured absorbancevalues at several wavelengths and potentially other measured valuesbased on absorbance versus wavelength (the X matrix) as processedthrough a kernel function as specified by the KOPLS method. Once theKOPLS coefficients of this relationship are established, they are usedwith the kernel function by the analyzer to predict the unknownhemoglobin parameter values from new measurements of absorbance onsamples.

The kernel function used in this example is a simple linear kernelfunction described in the Mattias Rantalainen et al. reference listedabove and represented by the following equation:

κ(X,X)=

X,X

where the matrix of measured values X is put into the kernel functionand subjected to further processing as specified in the cited KOPLSreferences above (incorporated by reference) for creating the KOPLStraining coefficients.

Once the set of training coefficients, or mapping function, isestablished, it is used to predict the hemoglobin parameter valuesand/or total bilirubin parameter values of a blood sample from futuremeasurements. A single-row X matrix is created from the newmeasurements, then the value from this single-row X matrix is putthrough the kernel and mapping functions to produce the hemoglobinparameter values and/or total bilirubin parameter values according tothe procedures necessary for the mapping function used according to theKOPLS procedures described in detail in the KOPLS references disclosedpreviously.

The data collected from the blood samples described above were putthrough the KOPLS method in a cross-validation process. Cross-validationis a process for using a data set to test a method. Several data rowsare set aside and the rest are used to create a mapping function. Theset-aside values are then used as “new” measurements and their Y matrixvalues calculated. This process is repeated by setting aside othermeasured values and computing another mapping function. By plotting theknown values of the blood data vs. the calculated, the effectiveness ofthe method may be ascertained by inspecting the plot.

Turning now to FIGS. 18-23, there are illustrated graphical plots of thecorrelation results comparing the various hemoglobin parameters of lysedblood to whole blood using the KOPLS method. The blood samples weremanipulated to cover a large range of analyte values. The technique ofn-fold cross-validation using 60 folds was used to test the data. Inthis technique, the data set is divided into n=60 separate sets, and themodel is made from n−1 of the sets, with the remaining set predictedusing the model. The process is repeated 60 times for each group. Everydata point is thus predicted using a model made from most of the otherdata points, without being included in the model.

FIG. 19 shows the correlation results for tHb using the K-OPLS method.The horizontal axis has units representing the total hemoglobin in gramsper deciliter of lysed blood. The vertical axis has units representingtotal hemoglobin in grams per deciliter of whole blood. As can be seenfrom the plot, the method of determining tHb of a whole blood sample hasa correlation of greater than 99%.

FIG. 20 shows the correlation results for O2Hb using the K-OPLS method.The horizontal axis has units representing the percent oxyhemoglobin oflysed blood. The vertical axis has unit representing percentoxyhemoglobin of whole blood. As seen from the plot, the method ofdetermining O2Hb of a whole blood sample has a correlation of greaterthan 99%.

FIG. 21 shows the correlation results for carboxyhemoglobin using theK-OPLS method. The horizontal axis has units representing the percentcarboxyhemoglobin of lysed blood. The vertical axis has unitrepresenting percent carboxyhemoglobin of whole blood. As seen from theplot, the method of determining COHb of a whole blood sample has acorrelation of greater than 99%.

FIG. 22 shows the correlation results for deoxyhemoglobin using theK-OPLS method. The horizontal axis has units representing the percentdeoxyhemoglobin of lysed blood. The vertical axis has unit representingpercent deoxyhemoglobin of whole blood. As seen from the plot, themethod of determining HHb of a whole blood sample has a correlation ofgreater than 99%.

FIG. 23 shows the correlation results for methemoglobin using the K-OPLSmethod. The horizontal axis has units representing the percentmethemoglobin of lysed blood. The vertical axis has unit representingpercent methemoglobin of whole blood. As seen from the plot, the methodof determining MetHb of a whole blood sample has a correlation ofgreater than 99%.

FIG. 24 shows the correlation results for tBil using the K-OPLS method.The horizontal axis has units representing the total bilirubin inmilligrams per deciliter of lysed blood. The vertical axis has unitsrepresenting total bilirubin in milligrams per deciliter of whole blood.As can be seen from the plot, the method of determining tBil of a wholeblood sample has a correlation of greater than 99%.

A method of making a whole blood measurement using the COOx analyzersubsystem 10 of the present invention will now be described. Anabsorbance scan is measured by first recording a transmitted lightintensity scan with cuvette module 43 filled with a transparent fluidsuch as water or analyzer flush solution otherwise known as the ‘blank’scan. Then a transmitted light intensity scan with cuvette module 43filled with the whole blood sample is recorded. After corrections forspectrometer dark response and detector linearity, the spectralabsorbance is the negative of the logarithm to the base ten of the ratioof the whole blood scan to the transparent fluid scan computed at eachwavelength in the measurement range.

More specifically, a depiction of the components of a COOx analyzersubsystem is shown in FIGS. 1-18. This subsystem embodiment measures theoptical absorbance of liquids introduced into cuvette module 43. Thelight used to perform the absorbance measurement originates from LEDlight source 28, is collected and transmitted by collimating lens 30,passes through first diffuser 32, circular polarizer 34, focusing lens36, and optional protective window 38 before reaching cuvette module 43.Critical to an absolute absorbance measurement is knowledge of thecuvette path length. The cuvette path length is pre-measured for eachindividual cuvette module 43 and programmed into an electronic chip 48 con cuvette module 43. The path length information is read/retrieved bydata processor module 130 of the analyzer whenever required.

After passing through cuvette module 43, the light is collected by lens66, collimated and sent through second diffuser 68 and beam splitter 69.The purpose of beam splitter 69 is to allow light from calibrating lightsource 72 (for example, a krypton gas-discharge lamp), collimated bylens 74, to enter optical path 21. Calibrating light source 72 provideslight at a few known wavelengths, which are used to periodicallyrecalibrate the wavelength scale of spectrometer module 100. Afterpassing through the beam splitter 69, the light is focused by lens 82onto an optical fiber 92. The optical fiber 92 guides the light to inputslit 114 of spectrometer module 100. The light passes through anachromatic lens 124, goes through light dispersion element 130 with areflective back 132. The light is wavelength-dispersed by passingthrough light dispersion element 130 such as, for example, prism 130then makes a return pass through the lens 124, which re-focuses thelight onto the pixels of light-array detector 116. Light-array detector116 converts the light energy into an electrical signal which representsthe spectral intensity of the light. The electrical signal is sent todata processor module 150 for further processing and display of thefinal results to the user. Light-receiving and converting assembly 110is a single board that holds input slit 114 and light-array detector 116in close proximity as an integrated unit.

Input slit 114 is applied directly onto the same circuit board substrate112 as and in close proximity to light-array detector 116. Other priorart spectrometers place these components on separate planes where theyhave separate mounting structures needing independent adjustment andalignment. The mounting scheme of the present invention has severaladvantages that lower the cost and size of spectrometer module 100: 1)cost of separate mounting structures is avoided, 2) input slit 114 canbe laser etched in a precise position relative to light-array detector116 making alignment less labor intensive, 3) inexpensive sphericalsurface optics can be used in the optical system since the image of theslit on the detector is only slightly off-axis from the center axis ofthe optical system, minimizing aberration, and 4) a single alignmentprocedure for a unified slit and detector assembly replaces alignmentprocedures for two separate assemblies.

It is important to note that first diffuser 32 and second diffuser 68are positioned before and after cuvette module 43, respectively. Opticalabsorbance measurement of a diffuse sample presents a unique problem.The diffuse transmittance of the sample scrambles the initial spatiallight distribution of the measurement system caused by the nonuniformitytypical of light sources. Thus, the spatial light distribution of the‘blank’ scan can be quite different from the whole blood sample scan.Since optical detectors have response that varies spatially, theresponse can vary due to spatial distribution changes of the incidentlight, even if the overall intensity has not changed. An absorbance scanwhich is based on the ratio of the sample scan to the blank scan willhave a significant absorbance component due to this effect in additionto the absorbance due to the sample alone. This results in a significantmeasurement error of the sample absorbance that is intolerable forcooximetry.

The advantage of placing cuvette module 43 between first and seconddiffusers 32, 68 is that the spatial light distribution will appear thesame for the blank and sample scans, removing this error effect.Diffusers 32, 68 are specially chosen so that they diffuse a ray ofincident light into the full acceptance cone of the optical system, butnot more so, so that as much light throughput as possible may bepreserved while scrambling the light ray completely across the field.

Although the preferred embodiments of the present invention have beendescribed herein, the above description is merely illustrative. Furthermodification of the invention herein disclosed will occur to thoseskilled in the respective arts and all such modifications are deemed tobe within the scope of the invention as defined by the appended claims.

What is claimed is:
 1. A replaceable cuvette assembly for use in anoptical absorbance measurement system for measuring whole-bloodhemoglobin parameters or whole-blood bilirubin parameters, the cuvetteassembly comprising: a cuvette substrate; and a cuvette module fixedlyconnected to the cuvette substrate wherein the cuvette substrate is asupport for securing the cuvette assembly within the optical absorbancemeasurement system, the cuvette module having a sample inlet port, asample outlet port, an electronic chip assembly, a sample receivingchamber that fluidly communicates with the sample inlet port and thesample outlet port, a first cuvette window, and a second cuvette windowforming a portion of the sample receiving chamber wherein the firstcuvette window and the second cuvette window are aligned with each otherdefining a cuvette optical path length between the first cuvette windowand the second cuvette window and disposed within an optical path of theoptical absorbance measurement system.
 2. The replaceable cuvetteassembly of claim 1 wherein the cuvette substrate has a cuvette lightpath opening through the cuvette substrate wherein the cuvette lightpath opening is disposed within optical path.
 3. The replaceable cuvetteassembly of claim 1 wherein the cuvette module includes a cuvette firstportion and a cuvette second portion wherein the cuvette first portionand the cuvette second portion are bonded to each other and forms thesample receiving recess.
 4. The replaceable cuvette assembly of claim 3further comprising a gasket disposed between the cuvette first portionand the cuvette second portion.
 5. The replaceable cuvette assembly ofclaim 1 wherein the defined cuvette optical path length is accuratelymeasured and stored within the electronic chip.
 6. The replaceablecuvette assembly claim 3 wherein the cuvette first portion contains asample receiving recess, the sample inlet port, the sample outlet port,the electronic chip assembly and the first cuvette window and whereinthe cuvette second portion contains the second cuvette window.
 7. Thereplaceable cuvette assembly of claim 3 wherein the cuvette firstportion further includes an electronic chip recess.
 8. The replaceablecuvette assembly of claim 3 wherein the second cuvette window of thecuvette second portion is a raised surface that forms a water-tight sealaround a sample chamber recess of the cuvette first portion.